1. Field of the Invention
This invention relates to the field of x-ray medical diagnostics, and more particularly to improving the transient response and stability in the control of diagnostic x-ray tube output in systems such as computerized tomographic equipment.
2. Description of The Prior Art
In use, a medical diagnostic x-ray tube is electrically actuated to produce x-rays which are directed through a patient's body. A pattern of x-rays which pass through the patient's body is sensed and the information obtained is used to produce a representation, most often a visual image, of internal body structure.
A typical x-ray tube includes a thermionic filament, or cathode, and an anode, both located within a substantially evacuated glass envelope. An electrical potential is applied across the filament. The resultant heating of the filament causes the filament material to emit, or "boil off", electrons, creating an electron cloud around the filament. A high potential is applied between the filament and an anode to accelerate electrons from the cloud so that they strike a target area of the anode and x-ray energy is emitted.
The flow of electrons from the cathode to the anode is known as the "anode current" or the "tube current", as distinguished from "filament current" which is the flow of electrons through the filament to effect its heating.
The rate of x-ray energy production is an increasing function of the tube current, other parameters being equal. For a fixed cathode to anode potential, tube current is an increasing function of the density of the electron cloud which in turn is a function of the potential applied across the filament. There is, therefore, a relationship between tube current and filament voltage. Typically the relation is exponential, i.e. nonlinear. Thus, the change in tube current which results from a given change in filament voltage is greater at higher filament voltages, i.e. higher tube currents, than for the same change at lower tube currents and filament voltages.
The life of an x-ray tube is a decreasing function of the output level at which it is operated, i.e. the intensity of x-ray energy it is stimulated to produce. Tube operating life is shorter at higher energy output levels than at lower operating levels.
Since the x-ray output is a function of the tube current, and tube current a function of filament voltage, it has been proposed to regulate x-ray output by controlling the tube current by regulating the filament voltage.
X-ray energy control is desirable because it enables optimum selection of a tradeoff in the tube current value selection wherein sufficient x-ray dosage is administered to achieve good imaging of the patient's internal body structure, while limiting the tube current and x-ray output sufficiently to enhance x-ray tube life.
In medical x-ray, the precision of tube current regulation required is a function of the type of study. In radiography, a relatively short, high intensity pulse of x-ray energy is directed through a patient's body, and a piece of x-ray sensitive film is exposed to create a "radiograph". In fluoroscopy, continuously pulsed or constant lower intensity x-ray energy is directed through the patient's body, from which it emerges to fall upon the input face of an image intensifier tube. The image intensifier tube converts the emergent x-ray pattern to a visible image at an output face which can be photographed, or viewed, as with television to observe changes within the patient's body.
In computerized tomography, (CT) a movable x-ray source is provided, along with an array of x-ray detectors. The x-ray source is moved about the patient's body and the x-rays from the source are directed through the body to the multiple detectors. Data processing equipment receives time varying signals describing information from the individual detectors and processes them to produce or "reconstruct" a tomographic image illustrating a planar segment taken through the patient's body.
In computerized tomography, the image is not reproduced directly in analog form as in the case of radiography and flouroscopy. Rather, the image is generated in response to complex time variations in electrical signals produced by each individual detector of the array.
In these radiographic and fluoroscopic systems the precision required in tube current control is not as great as in CT systems, in part because total doses are relatively low, and good imaging can be done over a relatively wide range of dose. More importantly, the x-ray film or fluoroscope time integrates all the energy falling upon it, and is not as sensitive to time variations in energy, but only to its total. In many radiographic devices, alternating current (AC) is used to heat the x-ray tube filament. The AC ripple appearing in filament voltage, and therefore in tube current, is not as significant in radiography, except in very short exposures, where ripple variation is not integrated on film, and in instances where the ripple causes large undesirable tube voltage changes. Also, the transient response of AC filament control is generally faster than for direct current (DC) control, due to required heavy filtering of transformer coupled filament drive circuitry used in A.C. application.
In CT scanning, the degree of precision required in x-ray output control is much greater than in radiography or fluoroscopy. This extreme precision in x-ray output stability is required because the time variations in detector outputs produced in response to detected radiation is crucial in enabling the computerized reconstruction of a quality image.
One means of enhancing precision control of x-ray output in CT is the use of direct current (DC) to heat the tube. DC is needed because the CT detectors would interpret AC ripple in detected x-rays as representing information about the patient's body.
Two general types of computerized tomographic scanners are the "translate-rotate" type (TR) and the "stationary detector" type (SD). In the TR type, the x-ray source is moved relatively slowly. In the SD type, in which an orbiting source is used in conjunction with a 360.degree. ring of stationary detectors positioned about the patient, the x-ray tube may move rapidly about a curved path around the patient's body.
The exercise of control for minimization of tube operating time is much more important in SD machines than in TR machines. In TR scanning, a stationary anode tube is used, and an oil bath is applied directly to the anode to cool it in order to counteract the effects of heating which results when the tube is driven at, e.g., about 6 kilowatts for scans requiring about 16-18 seconds. By dissipating the heat from the anode in the oil bath, the stationary anode tube can operate for relatively long periods at these fairly high output levels without unduly shortening its life. Moreover, the tube produces x-rays during each rotate portion of a scan so there is opportunity for cooling the anode when it is not producing x-rays.
A TR scanner also facilitates precise and periodic calibration of x-ray tube ouput. One can employ adjustment circuitry which can be operated to precisely calibrate the tube output and allow the tube to settle to a stable operating outputbefore actual data collection begins.
In the stationary detector machine, the need for even higher driving power (e.g. 28 kilowatts) makes a rotating anode type x-ray tube necessary and renders impractical the use of direct cooling of the anode by an oil bath.
Expressed another way, a stationary anode cannot be used (compared to radiographic work) at SD high power relatively long term loadings because its focal spot would overheat. Consequently, a rotating anode tube is used, because the rotating anode distributes the electron stream and the heat over a large anode section. It is not practical to directly oil cool a rotating anode, and therefore such a tube is more vulnerable to destructive heat buildup than the stationary anode variety having direct anode oil cooling.
In the stationary detector apparatus, it is therefore desirable to limit tube operating time to an absolute minimum so that life of the less efficiently cooled tube is not shortened more than necessary by heating. Consequently, the time allowed for calibrating the tube should be as short as possible, so that tube operation time is not extended any more than necessary outside the time of actual x-ray scanning duty in which data is collected. Nevertheless, the requirements of computerized tomography still demand that, during the actual scanning time, the tube output be as uniform and precisely controlled as possible. While the rise time of tube output should be as short as possible, overshoot in tube operating level from the steady state desired value should be minimal. To express this another way, the transient response of the x-ray tube output, upon actuation of the tube, should be approximately critically damped, the output rising in a smooth and rapid progression to the steady state level specified by the tube input parameters, there being minimal overshoot and subsequent oscillation of the tube output about the predetermined steady state value.
The problem of obtaining fast rise time in CT is aggravated by the need, as described above, for using DC filament current. The DC is inherently slower in control response than is the AC used in radiography.
Since the current output of many x-ray tubes is non-linear with respect to the controlling filament voltage, the transient response of the x-ray tube output current varies, dependent upon the steady state output current toward which the tube is driven upon actuation. This difference in transient response dictates that, in order for the transient response to be substantially uniform for various steady state currents selected, compensation must be performed with respect to the electrical circuitry actuating the tube, in order to insure that the transient rise to each selected steady state value be critically damped, neither suffering from the slow response of overdamped conditions, nor from the overshoot or instability associated with underdamping.
In practice, it has been found that it is desirable to be able to bring the tube operating level up to the steady state reference value within approximately 200 milliseconds (ms).
It is therefore an object of this invention to provide circuitry for controlling the tube current output of an x-ray tube to raise tube current to a predetermined desired steady state level, selected from a broad range of current values, in as short a time as possible consistant with the maintenance of stability in the output current, notwithstanding nonlinearity of the x-ray tube current output with respect to changes in filament voltage.